Embodiments of the invention relate generally to diagnostic imaging and, more particularly, to an organic x-ray detector assembly and a method for manufacturing an organic x-ray detector assembly.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. In particular, an x-ray tube included in the x-ray source emits the x-ray beam at a focal point. The beam, after being attenuated by the subject, impinges upon an array of radiation or x-ray detectors.
In known CT systems, the x-ray beam is projected from the x-ray source through a pre-patient collimator that defines the x-ray beam profile in the patient axis, or z-axis. The collimator typically includes an x-ray-absorbing material with an aperture therein for restricting the x-ray beam.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject so that the angle at which the x-ray beam intersects the subject is constantly changing. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a “view”. A “scan” of the subject comprises a set of views made at different gantry angles or view angles, during one revolution of the x-ray source and detector.
X-ray detectors for such CT imaging systems typically are configured in a circular arc centered to the focal spot. Such detectors include an additional collimator for collimating x-ray beams received at the detector with focus to the focal spot.
Conventional CT detectors also include a scintillator component and photodiode component adjacent the collimator. The scintillator component illuminates upon reception of radiographic energy and the photodiode component detects illumination of the scintillator component and provides an electrical signal as a function of the intensity of illumination. Each scintillator element of the scintillator component converts x-rays to light energy and discharges the light energy to an adjacent photodiode element. The light emitted by each scintillator element is a function of the number of x-rays that impinge upon the scintillator element as well as the energy level of the x-rays.
The photodiode component of typical CT detectors is manufactured using a rigid semiconductor material such as silicon. Each photodiode element in the CT detector detects the light energy and generates a corresponding electrical signal as a function of the light emitted by a corresponding photodiode element. The electrical signal generated by the photodiode element is indicative of the attenuated beam received by each scintillator element. The outputs of the photodiode elements are then transmitted to the data processing system for image reconstruction.
Each pixel in a generated x-ray image is formed based on the output signal from an individual photodiode element, which is fed to the image processing unit by way of a dedicated electrical channel bonded to the photodiode element. As such, high resolution image detectors (i.e., detectors with well over 10,000 pixels) include a complex pattern of electrical channels that run across the surface of the photodiode array or through internal layers within the photodiode array to electrically couple the respective photodiode elements to the digital readout electronics and/or application specific integrated circuits (ASICs). The portion of the surface of the detector that includes the electrical channels and bonding pads forms a dead zone on the detector surface. Electrode layers are affixed to contact points on the top and/or bottom sides of the semiconductor material to create a pattern of electrical.
High-resolution CT image detectors with silicon photodiodes are complex and expensive to manufacture and pattern due in part to the large number of conductor channels and connections between the photodiode elements and digital readout electronics. Further, precise alignment between the large number of respective pairs of rigid photodiode elements and scintillator elements further adds to manufacturing cost and complexity.
Therefore, it would be desirable to design a detector for a CT imaging system that overcomes the aforementioned drawbacks of conventional CT image detectors. It would further be desirable to reduce costs associated with fabricating a CT image detector.